Localized controlled release of therapeutics maximizes the effectiveness of a drug by selectively delivering the drug to a particular site. Local drug delivery decreases systemic exposure to a drug and avoids potentially dangerous peak and trough drug levels in patients. To this end, many devices have been developed to deliver drugs in a controlled, localized manner.
Although devices can be made into many different morphologies including microspheres, in-situ forming gels, films, and injectable depots, fibrous morphologies are intriguing because they allow for easy implantation/removal of a device, anchorability in tissue, and mechanical stability. (See, e.g., Langer, R., Chemical Engineering Science 1995, 50, (24), 4109-4121; Batycky, R. P., et al., Journal of Pharmaceutical Sciences 1997, 86, (12), 1464-1477; Freiberg, S., et al., International Journal of Pharmaceutics 2004, 282, (1-2), 1-18; and Wang, Y, et al., Biomaterials 2004, 25, (18), 4279-4285, the disclosures of each of with are incorporated herein by reference.) Fibrous devices can also provide mechanical stability or be woven with other fibrous components that are more mechanically stable.
As a result, encapsulating therapeutics within polymer matrices is a well established method of controlling the release of drug molecules. Biodegradable polymers are often used to control drug release and facilitate full device erosion. (See, e.g., Kohn, J., et al., Biodegradable and Bioerodible Materials. In Biomaterials Science Second Edition, Elsevier Academic Press: San Diego, Calif., 2004, pp 115-127; and Langer, R., 1995, the disclosures of which are incorporated herein by reference.) Copolymers made of lactide and glycolide monomers are common biodegradable polymers that are commercially available over a range of molecular weights and lactide to glycolide ratios. Since both molecular weight and lactide to glycolide ratio affect polymer degradation and therapeutic release, careful selection of a poly(lactide-co-glycolide) (PLGA) can achieve many temporally different release patterns. (See, e.g., Alexis, F., Polymer International 2005, 54, (1), 36-46; and Siepmann, J., et al., Advanced Drug Delivery Reviews 2001; 48(2-3):229-247, the disclosures of which are incorporated herein by reference.)
Common methods for making fibrous devices for drug delivery include electrospinning and wet spinning. (See, e.g., Kim, K., et al., Journal of Controlled Release 2004, 98, (1), 47-56; Kenawy, E. R., et al., Journal of Controlled Release 2002, 81, (1-2), 57-64; Zeng, J., et al., Journal of Controlled Release 2005, 105, (1-2), 43-51; Crow, B. B., et al., TISSUE ENGINEERING 2005, 11, (7/8); Crow, B. B., et al., Biopolymers 2006, 8, (6), 419-427; Gao, H., et al., Journal of Controlled Release 2007, 118, (3), 325-332; and Zilberman, M., et al., J Biomed Mater Res Part B: Appl. Biomater. 2005, 74, 792-799, the disclosure of each of which are incorporated herein by reference.) A particular method is usually chosen based on design criteria for the particular application and nature of the polymer. Electrospinning can be used to make fibers of micron to submicron diameters and is typically used to make non-woven meshes. (See, e.g., Fridrikh, S. V., et al, Physical Review Letters 2003, 90, (14), 144502; Subbiah, T., et al., Journal of Applied Polymer Science 2005, 96, (2), 557-569, the disclosures of which are incorporated herein by reference.)
Wet spinning is perhaps the earliest form of fiber processing and is done by dissolving a polymer in a solvent and extruding the solution into a non-solvent. The extrusion process typically happens by forcing the polymer solution through an orifice into a coagulation bath with an uptake bobbin pulling the fiber and spooling the final product. The solvent diffuses out of the polymer at rates dictated by miscibility of the solvent with the non-solvent, density of the polymer solution, temperature of the process, and other process parameters. The fiber coagulates when the polymer reaches a critical point and phase separates. Solvent is removed by prolonging the contact with the non-solvent, exposing the fiber to various non-solvent baths, or simply drying out the remaining solvent. The result is a solid polymer filament that may be further processed into yarn, textiles, sutures, or other devices. Traditional fiber drawing uses semi-crystalline polymers since the drawing procedure helps to align the crystalline regions of the polymer and enhance the mechanical properties in the direction of fiber drawing. To this end, post extrusion processing such as cold drawing is used to enhance crystallinity and strengthen fibers. Benefits of a wet-processed fiber include a porous microstructure, room temperature processing conditions, and diameters of suture-like scale. Fibers made from two phase solutions, such as emulsion or solid suspensions, have been shown to be effective in controlling therapeutic release and have been used primarily for tissue engineering and cancer therapy.
Poly(lactide) and poly(glycolide) are biodegradable polymers commonly used for medical applications. Their copolymer, PLGA, is commonly used for drug delivery applications. Poly(lactide) comes in two varieties, the single enantiomer poly(l-lactide) (PLLA) and the racemic poly(dl-lactide) (PDLLA). This slight difference in stereochemistry results in significantly different properties, since PLLA is semicrystalline and PDLLA is always amorphous. Poly(glycolide) is semicrystalline but copolymers of glycolide and lactide are almost always amorphous, the exceptions being copolymers with very high or very low glycolide to lactide ratios (l-lactide only). The semicrytalline versions of PLGA are useful for structural applications, such as sutures, since the crystallinity will increase the strength of the material. Amorphous PLGA is more useful for drug delivery, since the lack of crystallinity allows for the uniform distribution of drug throughout the polymer matrix. The release of drugs from PLGA can be tuned by the varying the ratio of lactide to glycolide since this ratio determines the degradation rate of the polymer. The fastest degrading PLGA is the 50:50 copolymer, with other copolymers taking longer to degrade as the ratio skews one way or the other. PLGA is a polyester and degrades by hydrolysis to give lactic acid and glycolic acid.
Drug release has been quantitatively linked to PLGA degradation by several groups, with one rigorous model described by Batycky et al. (See, e.g., Batycky, R. P., et al., Journal of Pharmaceutical Sciences 1997, the disclosure of which is incorporated herein by reference.) This model describes drug release from PLGA and PLA microspheres made by an emulsion process. This technique leaves isolated pockets containing protein throughout a polymer microsphere. As the spheres degrade, a system of interconnected pores forms between the pockets and the surface. The pores start at various sizes, but small pores (micropores) will not allow the protein to diffuse through. When large pores (mesopores) bridge between a certain pocket and the surface, the drug may diffuse out. This model takes into consideration hydration time for the polymer, degradation rates, and diffusivities and is able to accurately describe the drug release from the microspheres. These microspheres show a three phase release profile, with an initial burst, a period with low release, and a final period of sustained release. Hydration of the sphere should be quick (calculated at 8 min based on physical parameters of the spheres) so burst release is taken to be due to adsorbed molecules or drug that already is bridged to the surface by mesopores. Later work by found that at least some of the burst is due to the swelling of the pores associated with hydration. The swelling causes some of the pores to bridge to the surface of the sphere, causing rapid drug release. The second region of drug release is an induction period where little drug is released. This is due to initial porosity being low and most of the pores being small. The protein is unable to diffuse through the small, sparse pores and there is little drug release for a period of several days to weeks. The third phase of sustained release is due to polymer degradation causing the proliferation of mesopores, allowing the protein to escape. This three phase model of release is widely reported and is the main model of release for biodegradable polymer devices.
In a technical report, Nelson et al described a wet spinning procedure for PLGA and PLLA fibers. (See, Nelson, K. D., et al., Tissue Engineering 2003, 9, (6), 1323-1330, the disclosure of which is incorporated herein by reference.) This report focuses on semi-crystalline PLLA fibers but mentions that PLGA fibers made by the same techniques degrade faster than PLLA fibers (2 months for PLGA compared to more than 5 months for PLLA). These fibers are mechanically stable enough for a tissue engineering scaffold with PLLA being particularly useful due to the longer time for degradation. This technique was used to encapsulate drug by an emulsion method. (See, Crow, B. B., et al., TISSUE ENGINEERING 2005, 11, (7/8), the disclosure of which is incorporated herein by reference.) Protein was dissolved in water, which was emulsified in a polymer/methylene chloride solution. The emulsion was extruded into hexanes for coagulation with a draw ratio of 41. The fibers in this case had a maximum drug loading of 2.38 wt % to start, but final drug loadings are not provided. The process for making the fibers is very similar to processes for making w/o/w emulsion based microspheres, and the release profile is also very similar. There is an initial burst of drug ranging from 5-15%, followed by an induction phase lasting to 5 weeks, and a release portion governed by degradation. The burst region is higher for PLLA fibers and for the PLLA fibers, fibers with 10% aqueous loading have a higher burst than 5% aqueous loading. Oddly, the induction phase seems to be equivalent for PLLA and PLGA despite the much faster degradation of PLGA. In the degradation release phase, PLGA releases much faster than PLLA. It is unclear why PLGA releases the vast majority of the drug load from 5-11 weeks when complete polymer degradation happens at 7 weeks.
Building on the emulsion fibers, Gao et al used a simple particle suspension instead of an emulsion to control the release of 5-fluorouracil from PLLA fibers. (See, e.g., Gao, H., et al., Journal of Controlled Release 2007, 118, (3), 325-332, the disclosure of which is incorporated herein by reference.) The spinning technique did not use any uptake, so diameters were varied and depended upon the molecular weight of the polymer and polymer concentration. No diameter attenuation was used. The result was a fiber with high drug loading (75-90% of initial drug) and diameters spanning from 65-152 mm. Drug release was monitored only through the burst phase, with maximum measurements at 25 days (corresponding to less than a 1% drop in fiber mass). In this way, drug release was not dependent on polymer degradation but was instead controlled by thickening the polymer phase and shrinking the particle size of the drug. Most of the fibers described had sudden bursts of 10-90% followed by induction phases, but sometimes the burst could be slowed to give a controlled release over several days. Mechanical properties were not examined, but work by Zilberman et al on wet-spun PLLA fibers for stents suggests that these fibers should retain almost total strength over at least 6 weeks. (See, e.g., Zilberman, et al., J Biomed Mater Res Part B: Appl. Biomater. 2005, the disclosure of which is incorporated herein by reference.)
Although these techniques and experiments have shown that there is promise in these fiber release systems, a drawback to these systems is that drug release due to degradation cannot be decoupled from mechanical properties unless release is diffusion limited. This limits the usefulness of such devices since degradation controlled release is desirable for many circumstances, including long term release and modulated release. In addition, by using only a single polymer, release can only occur over a single given time interval. Moreover, these devices can only encapsulate a single drug and conditions for encapsulation and controlled release may not be general across all drugs. Accordingly, there is a need for improved fibrous drug delivery systems capable of having a more flexible drug release mechanism.